This invention relates to ocular prostheses, and in particular to prosthetic replacements for the eyeball. In one embodiment, the invention provides an eyeball prosthesis of an improved material.
It is common for an eye to have to be removed because of severe trauma, infection, congenital abnormality, untreatable painful glaucoma, or the presence of a tumour. After removal of an eye it is desirable to place a prosthetic globe of similar volume within the eye socket cavity. If this is not done in children the orbit orbital bones fail to develop normally, making wearing of a cosmetic shell difficult, and resulting in an unsightly appearance (Soll, 1982; Kennedy, 1973; Berry, 1991). In an adult, although volume replacement is not necessary in terms of bone development, which is already complete, it is vital for orbital volume to be maintained if the patient is to be fitted with an external cosmetic shell and to obtain a natural appearance and xe2x80x9ceyexe2x80x9d movement. A prosthesis is required for similar reasons if an eye is congenitally absent. Thus there is a widespread need for eyeball prostheses (socket prostheses).
The usual procedure on removal of an eye is to dissect off and preserve the covering conjunctiva, Tenon""s capsule (a thicker connective tissue layer which forms the fascial sheath of the eyeball) and the muscles responsible for eye movement. A prosthetic globe is then pushed into the space, the muscles are attached to the prosthesis, and covering tissues are sutured over the outside (Nunery et al, 1993). Once the tissues have healed, an external cosmetic shell can be worn over this surface under the eyelids, and should give reasonable cosmesis and some movement, as it is xe2x80x9ccarriedxe2x80x9d by movement of the implanted prosthesis. However, although,in principle an effective prosthesis is possible, major technical difficulties remain to be overcome.
The criteria for an ideal ocular prosthesis are simple; it should be buried in the existing eye socket, and it should be simple, light, smooth and inert (Soll, 1982). In particular, such an ideal prosthesis will not provoke any inflammatory response, and will permit suturing of the ocular muscles directly to the prosthesis so as to provide movement simulating that of the patient""s own eye.
It is clear from the literature that the ideal orbital implant has not yet been achieved. Mules appears to have been the first surgeon to place a hollow glass sphere, which he called xe2x80x9cartificial vitreousxe2x80x9d, within the scleral cavity of an eviscerated eye in human patients with the aim of improving cosmesis; infection-related extrusion was the greatest problem experienced (Mules, 1884). Lang and Frost both developed this idea by using a glass ball as a volume replacement after enucleation; Lang subsequently used celluloid balls to avoid the risk of breakage of a glass prosthesis (Lang, 1887).
Jardon (U.S. Pat. No. 2,688,139) proposed a ball with irregular surface and perforations to encourage tissue ingrowth. Polytetrafluoroethylene, polyethylene, poly(methyl methacrylate) and nylon were suggested as materials.
In recent years simple poly(methyl methacrylate) or silicone spheres have become the norm, in spite of extrusion rates of over 10% in some series (Nunery, 1993). However, it is impossible to attach muscles directly to these materials. Therefore the implant must first be covered with donor cadaveric sclera or with a synthetic material such as Dacron (trade mark), and the muscles are then sutured to the covering layer.
Recently prostheses made of hydroxyapatite have been used in some centres (Perry, U.S. Pat. No. 4,976,731; Dutton, 1991), and these are thought to give a better outcome because, since hydroxyapatite is porous, the surrounding tissues can grow into the ball and help to secure it. However, hydroxyapatite is a hard material and cannot be sutured directly to the tissues; therefore it must still be covered with sclera or Dacron. Furthermore, hydroxyapatite implants are extremely expensive (Dutton, 1991; McNab, 1995) and are thus not routinely available to patients even in many economically developed countries.
Hydroxyapatite implants have gained in popularity in those countries where they are affordable, and several large series have been reported (Dutton, 1991; Shields et al, 1993; Shields, 1992; Shields, 1994). There is evidence that fibrovascular ingrowth into hydroxyapatite does take place, although it appears to be accompanied by a degree of inflammation of the foreign body type (McNab, 1995; Shields et al, 1991; Rosner et al, 1992; Rubin et al, 1994). Moreover, extrusion of the prosthesis can still occur (McNab, 1995; Buettner and Bartley, 1992).
Rubin et al (1994) reported a study in rabbits, in which spheres of hydroxyapatite, 14 mm in diameter, were compared with spheres of porous polyethylene. Vascularisation occurred more quickly in the hydroxyapatite spheres, especially near to the muscle insertions, but both materials showed a low-grade foreign body response.
Girard (1990) has proposed Proplast (trade mark) as both an enucleation implant, citing Neuhaus (1984), and as an evisceration implant, and describes his evisceration technique in the pig in 4 eyes with a maximum of 1.5 years follow-up. This work appears promising, but the results of more extensive trials are awaited. Some studies of keratoprosthesis development have found Proplast to be prone to an inflammatory response and resultant extrusion from the eye (Legeais et al, 1992).
Vachet (U.S. Pat. No. 5,089,021) proposed a ball silicone elastomer covered with porous polytetrafluoroethylene. The porous material is not incorporated into the ball, and is attached either by suturing (like covering a wooden cricket ball with leather) or by adhesives.
Jacob-LaBarre (U.S. Pat. No. 5,192,315) proposed a silicone ball covered with xe2x80x9cpatchesxe2x80x9d of porous silicone (for cellular ingrowth) and other porous polymers (for muscle attachment) such as polyurethanes, polyesters, or polytetrafluoroethylene. The means by which the patches are adjoined to the underlying prostheses is not explained. Bare areas between patches are said to enhance mobility by preventing too much tissue attachment.
Goldberg et al (1994) have proposed using porous polyethylene implants which, like Proplast, do not require wrapping, and which allow muscles to be sutured directly to the implant. They reported trying the material in 16 rabbits; there were 2 early extrusions related to infection. Hydroxyapatite spheres used for comparison in 2 rabbits seemed to provoke more inflammation, but fibrous encapsulation occurred with both. The polyethylene showed fibrovascular ingrowth and macrophage invasion; however, the fibrovascular tissue did not reach the centre of the sphere.
De-epithelialized dermal fat grafts have also been used as primary as well as secondary orbital implants (Smith and Petrelli, 1978; Migliori and Putterman, 1991; Smith et al, 1988; Borodic et al, 1989); however, resorption of the fat may occur. Buccal mucous membrane grafts are of use in the contracted socket as a secondary procedure, providing both volume and mucosal surface (Molget et al, 1993). Unless autografts are employed, such fat or mucous membrane graft would entail a risk of graft rejection, as well as a high risk of disease transmission.
Currently the globe prostheses used are generally made of silicone (Nunery et al, 1993; Soll, 1974; Nunery, 1993) in a range of sizes. These prostheses themselves cannot integrate with tissues, or have muscles directly attached to them. They therefore have to be covered before implantation with either donor sclera from a cadaver (Soll, 1974), or with Dacron (trade mark) mesh. Dacron is expensive, and often fails to prevent extrusion, possibly because it does not cushion the prosthesis in any way; if the covering tissues are thin or weak the prosthesis may erode through them. Once such erosion has occurred re-implantation is technically very difficult, and cosmetic appearance and movement are often unsatisfactory. The use of cadaver tissue is undesirable because of the high risk of disease transmission and because of the chronic shortage of donor tissue.
Thus it can be seen that the prosthetic materials and techniques available in the prior art still suffer from problems of:
infection;
inflammation due to foreign body response;
poor attachment of extraocular muscles;
poor integration with surrounding tissues;
erosion of covering tissues;
extrusion of the prosthesis;
expense of certain materials; and
high risk of transmission of disease such as HIV, hepatitis virus and Creutzfeld-Jacob disease when transplanted cadaver tissues are used.
Previous work carried out in the applicant""s laboratory has shown that a particular hydrophilic polymer has considerable advantages in the manufacture of a composite corneal prosthesis (keratoprosthesis). The polymerisation characteristics of this polymer enabled production of a prosthesis having an annular opaque, spongy peripheral zone surrounding a central, transparent optic zone, in which the two zones were joined by an interpenetrating polymer network. The surrounding tissues were able to invade the spongy peripheral zone following implantation (Australian Patent No. 650156, equivalent to U.S. Pat. No. 5,458,819, and Chirila et al, 1994). Because of the functional requirements of the cornea, it was essential that the polymer used for the centre of the keratoprosthesis be capable of being transparent.
We have now surprisingly found that a hydrogel comprising a biocompatible hydrophilic polymer such as poly(2-hydroxyethyl methacrylate) can be used to form a socket prosthesis of novel construction. The physical and chemical properties of this material provide a number of significant advantages.
According to a first aspect, the invention provides an ocular socket prosthesis comprising a hydrogel consisting essentially of a biocompatable hydrophilic polymer on to which tissues can be directly sutured.
Preferably the prosthesis comprises the polymer both in its homogeneous gel form and in its sponge form, and the two forms are chemically joined at their interface via an interpenetrating polymer network (IPN). However, it is also contemplated that the prosthesis of the invention may be made predominantly or entirely from the sponge form of the polymer.
In one preferred embodiment, the prosthesis is generally spherical, and comprises a posterior hemisphere consisting essentially of the gel form of the polymer and an anterior hemisphere consisting essentially of the sponge form of the polymer, said anterior and posterior hemispheres being permanently joined at their interface by an IPN.
Optionally the anterior hemisphere is strengthened by a column of the gel form of the polymer, which extends from the surface of the anterior hemisphere towards the centre of the prosthesis; the gel column may extend to meet the posterior hemisphere, but need not do so.
In an alternative preferred embodiment, the prosthesis is generally spherical, and comprises two generally concentric regions, with a generally spherical central core consisting essentially of the homogeneous gel form of the polymer surrounded by a layer of the sponge form of the polymer, the core and the outer layer being permanently joined by an IPN.
It will be clearly understood that other means of attachment between the gel and sponge components of the prosthesis are within the scope of the invention; for example, a variety of suitable adhesives acceptable for surgical use will be known to the person skilled in the art. However, it is considered that the permanent joining of the preferred embodiments of the invention provides a significant advantage.
In both embodiments, the anterior surface of the prosthesis may optionally carry connecting means adapted to cooperate with corresponding receiving means on an externally-worn cosmetic shell. For example the prosthesis may have an anterior point, nipple, hole or groove adapted to interlock with a corresponding means in the cosmetic shell. It is considered that this would be safer than breaching the conjunctival covering tissue by placing a peg into the prosthesis in order to couple the prosthesis to the shell, as has been used in conjunction with hydroxyapatite implants (Shields et al, 1993).
In the prosthesis of the invention, the solid gel form of the polymer and the sponge form of the polymer are permanently chemically joined at their interface via a tight attachment which is achieved by the penetration of the hydrophilic monomer solution into the pores of the polymer sponge. Thus a sequential IPN is formed by penetration of the monomer through diffusion into the substance of the sponge, ie. by swelling. To some extent the solid gel is polymerised on the matrices of the sponge. Thus the prosthesis is produced by a sequential two-stage polymerisation performed in the same mould, in which the hydrogel sponge portion is formed first; during the second stage the monomer mixture firstly penetrates to some extent into the pre-existing polymer matrix and subsequently undergoes polymerisation, so as to form an interpenetrating polymer network along the boundary between the two portions.
Any biocompatible hydrophilic polymer capable of forming both a gel form and a spongy form of the polymer will be suitable for use in the invention. However, we have found that poly(2-hydroxyethyl methacrylate) (PHEMA) is particularly suitable for the purposes of the invention. The monomer, 2-hydroxvethyl methacrylate (HEMA), would normally be used at a concentration of 5 to 25% by weight in water. A cross-linking agent (0.2% to 1% by weight, based on the total amount of monomer) and a water-soluble initiator (0.05% to 1% by weight, based on the total amount of monomer) is added in order to effect polymerisation.
Optionally other monomers may be used in conjunction with HEMA to form a copolymer. These may be hydrophilic or hydrophobic monomers, and include for example other hydroxylated methacrylates and acrylates, acrylamide derivatives and N-vinylpyrrolidone and combinations thereof; hydrophobic methacrylates and acrylates; and other hydrophobic monomers. The person skilled in the art would be aware of many suitable comonomers. For example, a number of suitable compounds are described in our earlier Australian Patent No. 650156. This earlier patent also discusses in detail methods for production of the polymers. However, the person skilled in the art will appreciate that for the purposes of the present invention it is not necessary that any part of the prosthesis be transparent.
In selecting an appropriate material for the prosthesis of the invention, it must be borne in mind that the prosthesis must be capable of being sterilised, preferably by autoclaving at 120xc2x0 for 20 minutes. The prosthesis will normally be stored in sterile solution, such as phosphate buffered saline, until implantation.
The prosthesis of the invention may optionally be impregnated with collagen prior to implantation, in order to stimulate cellular ingrowth from the surrounding tissues. For example, the sterilised prosthesis is incubated for 18 to 24 hours at 4xc2x0 in an aqueous solution of sterile collagen (1.35 mg/ml) at pH 7.4, then warmed to 37xc2x0 and incubated for 1 hour before being transferred into sterile phosphate buffered saline for storage prior to implantation.
In order to reduce the risk of infection, the prosthesis may also be soaked in antibiotic solution before implantation.
Other optional treatments before implantation, for example to modify the tissue response, are also within the scope of the invention.